1. Field of the Invention
The present invention generally relates to magnetic resonance imaging coils and, more particularly, to z-gradient shielding coils.
2. Background of the Invention
Magnetic resonance imaging (MRI) systems are currently employed in forming images of the internal human anatomy. In such systems, a patient is placed in a magnetic field and is subjected to radio-frequency electromagnetic pulses. The magnetic resonance of the atomic nuclei of the patient are detected with a radio frequency receiver to provide information from which an image of that portion of the patient containing these nuclei may be formed. The magnetic field includes a main magnetic field and three additional fields with linear spatial gradients in the x, y, and z directions
The main magnetic field is a very strong magnetic field, which may be created by a super-conducting coil, a resistive coil, or a permanent magnet. Normally, the z-axis is parallel to the axis of the main magnetic field for systems in which the magnet has cylindrical geometry, such as for whole body imaging. The linear gradient magnetic fields are typically created by resistive coils and are referred to as gradient coils. The resistive coils create a magnetic field within the coil with a linear spatial gradient, also referred to as a magnetic gradient. Typically, there is one gradient coil for each of the x, y and z-axes, which create x, y, and z magnetic gradients, respectively. Two different types of gradient coils are typically used to produce the magnetic gradients for MRI, one which creates a magnetic gradient along the z (or longitudinal) axis of the coil, and two others which create magnetic gradients along either the x or y (transverse) axes.
In operation, for imaging purposes, it is necessary to rapidly pulse electrical current through the three gradient coils. When this is done, a problem commonly encountered is the induction of eddy currents in various metallic parts of the MRI system. The MRI system typically contains a metallic cylinder called a bore tube. The inside of the bore tube is an image volume; however, most imaging occurs only in the central portion ofthe bore tube. The current in the gradient coils induce eddy currents in the bore tube of the MRI system that, in turn, induce a magnetic field within the image volume, referred to as the error field. The magnetic field created by the eddy currents is undesirable in the image volume. In many medically useful imaging procedures, it is highly desirable to reduce or eliminate these eddy currents.
Typically, eddy currents are reduced by surrounding each gradient coil, also referred to as the inner coil, with another similar coil, an outer coil or a shielding coil, to cancel the magnetic and induced electric fields in the region outside of the outer coil. A set of a gradient coil and its associated shielding coil is referred to as a shielded gradient coil set. Ideally, the shielding coil is designed to exactly cancel the electric and magnetic field outside of the coil set. If no field exists outside of the shielded gradient coil set, then no eddy currents can be induced in the metallic parts of the MRI system, and therefore, no error field will be produced in the image volume.
Not all eddy currents affect the imaging volume equally, in particular, the induced eddy currents and magnetic fields can be analyzed in terms of the azimuthal harmonic number, m. The azimuthal harmonic number m means that the field or gradient varies in azimuthal angle like cosine(mxcfx86), sin(mxcfx86), or a linear combination of the two, where xcfx86 is the azimuthal angle as shown in FIG. 5B. That is, the field goes through exactly m full cycles as the angle varies from 0 to 360 degrees. The worst effects are seen from eddy currents with m=0. These harmonics also have the longest lifetime, which can be as long as several seconds. In general, the lower the m number, the worse the effects on an MRI system.
Existing attempts to reduce the eddy current effects have only been partially successful, especially for the z-gradient coil set. One common technique for making z-gradient coils is using circular parallel loops of wire, all of which lie in planes that are perpendicular to the z-axis. The loops are interconnected by straight wires that lie on the outer cylindrical surface of a support structure and are parallel with the z-axis. This design has the advantage that it creates no x or y magnetic gradient. This is important because it is undesirable to use a z-gradient coil that creates x or y magnetic gradients. However, the z-gradient that is created is not exactly homogenous, but varies with the radius from the z-axis.
The problem with this conventional shielding is that it is impossible to exactly cancel the field outside of the z-gradient coil set. A continuous surface current distribution would be required on the surface of the shielding coil to exactly cancel the field outside of the gradient coil set. Conventional shielding simulates a continuous surface current distribution by winding several discrete circular loops around a support structure. However, these discrete circular loops cannot exactly simulate a continuous surface distribution, and therefore, never exactly cancel the field outside of the gradient coil set.
While is not possible to exactly cancel the entire field outside of the gradient coil set, it would be desirable to cancel the specific harmonics that are most troublesome to the MRI system. Therefore, a shielding coil for a z-gradient coil that exactly cancels the magnetic fields of low azimuthal harmonic number, m, outside of the z-gradient coil set would be very desirable.
FIG. 1 illustrates an exemplary prior art MRJ system 10 as disclosed in U.S. Pat. No. 4,733,189. As shown in FIG. 1, the MRI system 10 includes a main magnetic component 20, gradient coils 30, shielding coils 40, and a detection component 50.
The main magnetic component 20 can be a permanent magnet, a resistive electromagnet, or a superconducting system as shown, in which a solenoidal electromagnet 22 is encased within a cryogenic vessel 26. Bore tube 28 supports the solenoidal electromagnet 22. Image volume 24 is located centrally to the main magnetic component 20.
Gradient coils 30 include an x-gradient coil 32, a y-gradient coil 34 and a z-gradient coil 36, disposed to create gradient fields orthogonal to each other. X and y gradient-producing coils are preferably implemented by saddle-shaped coil elements disposed about the main magnetic field axis and rotated ninety degrees from each other in orientation. As shown, the z-gradient coil 36 is implemented by a parallel loop gradient coil coaxial with the main magnetic field axis.
Detection component 50 includes a radio frequency (RF) coil 52 and an RF interrogator 56 and receiver 58. The interrogator 56 produces a pulse of radio frequency excitation and the energy emitted as the atoms return to an aligned state is captured via coil 52 and used to obtain an image signal. In use, a patient or other object is positioned within the image volume 24 of the system 10.
Shield component 40 includes an x-shielding coil 42, a y-shielding coil 44, and a z-shielding coil 46 disposed to counteract the eddy currents induced by the gradient-producing coils 32, 34 and 36, respectively. The x and y shielding coils, 42 and 44 may be implemented by saddle-shaped coils cut from flat copper sheets and rolled into the appropriate saddle shapes. As shown, the z-shielding coil 46 is implemented by a parallel loop shielding coil coaxial with the main magnetic field axis.
FIG. 2 illustrates an exemplary prior art parallel loop gradient coil of a type which may be used as a z-gradient coil 36 in FIG. 1. As shown in FIG. 2, parallel loop gradient coil 80 includes loops 81 interconnected by straight wires 82 that lie on the outer cylindrical surface of a support structure 84 and parallel with the z-axis. The loops 81 and straight wires 82 are formed from a single wire 86. Terminal connections 88 are connected to both ends of the single wire 86. This design has the advantage that it creates no x or y magnetic gradient. This is important because it is undesirable to use a z-gradient coil that creates x or y magnetic gradients. However, the z-gradient that is created is not exactly homogenous but varies with the radius from the z-axis. The number of loops is determined by the current available and the gradient desired. Typically, two pairs of loops are used, called a Maxwell pair.
FIG. 3 illustrates an exemplary prior art multiple winding gradient coil as described in U.S. Pat. No. 5,289,129 to Joseph, which may be used as a z-gradient coil 36 in the exemplary MRI system of FIG. 1. As shown in FIG. 3, multiple winding gradient coil 100 includes two windings. The first electrically conductive winding 112 is wound about the surface of cylindrical coil support structure 110. The first electrically conductive winding 112 is wound helically in a synmmetric manner about the center 114 of the coil support structure 110. Terminal connections 116 are connected to the first winding 112. A second electrically conductive winding 118 is also wound in an interleaved manner with respect to the first winding 112. In accordance with the invention, the second winding 118 is offset in azimuthal angle by 180xc2x0 (360xc2x0/2 windings) with respect to the first winding 112. For ease of illustration, the wire diameter of the second winding 118 has been illustrated to have a smaller diameter than the wire diameter of first winding 112. As with the first winding 112, second winding 118 also has a terminal connection 120 to which current is applied from a power supply (not shown) for generating a magnetic field. In another embodiment, the winding gradient coil may include a plurality of windings, offset in azimuthal angle by 360xc2x0/X windings. Any number of windings can be used, but for simplicity, only two are shown. The contents of U.S. Pat. No. 5,289,129, are hereby incorporated by reference for ease of description.
FIG. 4 illustrates an exemplary prior art parallel loop shielding coil, for use as a shielding coil 46 as shown in FIG. 1. As shown in FIG. 4, loops 130 are interconnected by straight wires 132 that lie on the outer cylindrical surface of a support structure 134 and are parallel with the z-axis. The loops 130 and straight wires 132 are formed from a single wire 136. Terminal connections 138 are connected to both ends of the single wire 136. Experience indicates that existing art is only partially successful in reducing eddy current effects, especially for the z gradient coil set.
A z-gradient shielding coil and a z-gradient coil set is desired that improves upon the coils and coil sets of the prior art to reduce eddy currents induced in the MRI system, particularly eddy currents with low azimuthal harmonic number, m. The present invention has been developed to address these needs in the art.
The above mentioned needs are met by a cylindrical whole body magnetic resonance imaging system gradient shielding coil having multiple windings which are individually wound about a cylindrical support structure such that the windings are evenly distributed in azimuthal angle and interleaved with one another. The windings are preferably spaced in azimuthal angle by 360xc2x0/N, where N is the number of windings. In such a design, m azimuthal harmonic components of the error field caused by the effects of eddy currents within the imaging volume can be canceled. Canceling an increasing number of m components requires increasing the N number of windings.
The cylindrical shielding coil of the present invention provides a canceling magnetic field gradient to cancel an error magnetic field gradient created within the imaging volume by eddy currents generated by a gradient coil in a volume outside of the gradient coil. The shielding coil includes a non-magnetic electrically insulating cylindrical coil support having an internal cavity which forms a volume for accepting the gradient coil and the imaging volume. N electrically conductive cylindrical windings are wound in a plurality of turns in a substantially helical path about a surface of the coil support, each turn of each winding being electrically spaced from each other turn of each the winding such that spacing between respective turns of each the winding decreases in approximate proportion to the distance of the respective turns from a center of each of the windings in a direction parallel to the axis of the coil. Each of the N windings is interleavingly wound in the same direction about the surface of the coil support and separated from each other winding of the N windings in an angular orientation of approximately 360xc2x0/N about the coil support. N is determined so as to cancel all harmonics up to and including the Mth harmonic of the error magnetic field gradient.
The present invention also includes a gradient coil set for use in a magnetic resonance imaging system. The gradient coil set includes a cylindrical gradient coil and a cylindrical shielding coil. The shielding coil is as described above. The gradient coil may be any cylindrical gradient coil.
The present invention also includes a gradient coil set electrically connected together to provide a fraction of the current to the shielding coil. The gradient coil may include a plurality of gradient windings with an inductance and a resistance. The shielding coil may include a plurality of shield windings with an inductance and a resistance. The gradient coil set may include at least one external coil having an inductance and resistance substantially equivalent to the inductance and resistance of the gradient winding of the shielding coil. The plurality of shield windings are connected in parallel. One of the plurality of gradient windings are connected in series to the plurality of shield windings. The remaining gradient windings are connected in series with an external coil. The combination of the plurality of gradient windings connected in series to the plurality of shield windings is connect in parallel to the combination of the remaining gradient windings connected in series with the external coil.